Previously, left ventricular assist devices have been commonly used to assist the heart pump blood. Current generations of the left ventricular assist devices included rotary blood pumps that were a significant improvement beyond the previous versions of pulsatile pumps, which relied on compression pusher plates or pneumatic drive systems cooperating with one-way valves to provide a pulsatile blood flow. The main advantage for rotary blood pumps is that they are typically considerably smaller and include less points or areas of stagnant blood flow, and thereby may reduce or avoid the risk of haemolysis or the formation of blood clots.
The rotary blood pumps typically provide a relatively continuous flow output by way of a centrifugal or axial type impeller, which is electrically urged to rotate within a housing to provide the pumping force. These continuous flow pumps generally are connected in parallel with the heart of a patient. This parallel flow is typically made possible by connecting the rotary blood pump to the apex of the ventricle through a cored hole and then pumping the blood to a cannulated area of the aorta or pulmonary artery.
One of the main disadvantages of these rotary blood pumps is that the output flow is relatively continuous and the output of the natural heart is pulsatile thus it is difficult to control pressures within the heart using rotary pumps. Furthermore, rotary pumps generally do not intrinsically change their pumping activity sufficiently in response to normal changes in blood flow. In some instances where inflow to the heart is transiently reduced, blood is effectively being sucked out of the ventricle at a rate faster than the filling rate of the ventricle. This generally results in the ventricle suffering a suction event whereby the interventricular septum and free wall are pulled together. This pulling together of the septum and the ventricle wall may partially or fully occlude the inflow cannula of the rotary blood pump, thereby stopping or limiting flow from both the natural heart and the rotary blood pump. This reduction in ventricular pressure can also cause collapse of the heart's atria or the veins supplying the heart, similarly reducing the amount of blood able to be pumped. In addition, gross disturbance of the shape of the ventricle can lead to dangerous cardiac arrhythmias. Thus, these suction events preferably should be avoided.
Furthermore, in cases of diastolic heart failure, stiffening of the ventricular wall may impede the filling of the heart and thus require higher effective filling pressures to achieve normal ventricular volumes. In this condition, even marginal mismatches between the amount of blood being pumped by the blood pump and the amount arriving at the heart can cause dangerous rises in venous pressure resulting in life-threatening congestion of either the lungs or systemic veins.
An example of a rotary blood pump is described within U.S. Pat. No. 6,227,797—Watterson et al. This rotary blood pump is a generally centrifugal device utilising a hydrodynamic suspension to levitate a magnetically rotated impeller.
Previously, U.S. Pat. No. 6,945,998—Liotta et al described a pulsatile blood pump connected between the left atrium of a heart and the aorta. However, the blood pump described in this specification is a pulsatile pump, which includes two mechanical valves and a system of pusher plates. The combined result of these features is that there are multiple possible stagnation points for blood clots to be formed within the described pulsatile pump.
Furthermore, pulsatile pumps are generally incapable of generating significant suction events in the natural heart. Rotary blood pumps work in a mechanically different manner and generally may induce suction events.
The present invention aims to or at least address or ameliorate one or more of the disadvantages associated with the abovementioned prior art.